Opposite to conventional radiography wherein an intensifying luminescent phosphor screen directly emits luminescent radiation and wherein said screen is not a storage medium, radiation image recording and reproducing techniques utilizing a radiation image storage panel, referred to as the stimulable phosphor screen, sheet or panel, are provided with a stimulable phosphor. With radiation image recording and reproducing techniques, the stimulable phosphor of the radiation image storage panel is caused to absorb radiation, which carries image information of an object or which has been radiated out from a sample. Said stimulable phosphor is exposed to stimulating rays, such as visible light or infrared rays, which causes the stimulable phosphor to emit light in proportion to the amount of energy stored thereon during its irradiation exposure.
The emitted fluorescent light is then photoelectrically detected in order to obtain an electric signal. The electric signal is further processed, and the processed electric signal is utilized for reproducing a visible image on a recording material. This way of working, making use of storage phosphor sheets or panels as an intermediate storage medium is also called “computed radiography”.
As in radiography it is important to have excellent image quality for the radiologist to make an accurate evaluation of a patient's condition, important image quality aspects are image resolution and image signal-to-noise ratio.
For computed radiography signal-to-noise ratio depends on a number of factors.
First, the number of X-ray quanta absorbed by the storage phosphor screen is important. Signal-to-noise ratio will be proportional to the square-root of the number of absorbed quanta.
Second, the so-called fluorescence noise is important. This noise is caused by the fact that the number of photostimulated light (PSL) quanta detected for an absorbed X-ray quantum is small. Since a lot of the PSL light is lost in the detection process in computer radiography, fluorescence noise has an important contribution to the signal-to-noise ratio. It is important that, on the average, at least 1 photon is detected for every absorbed X-ray quantum. If this is not the case, many absorbed X-ray quanta will not contribute to the image and signal-to-noise ratio will be very poor.
This situation is most critical in mammography, where X-ray quanta are used with low energy. Softer X-ray will give rise to less PSL centres and, therefore, to less PSL photons than harder X-rays.
In computer radiography, a number of PSL centres are created by the absorbed X-ray quanta. Not all PSL centres are stimulated in the read-out process, because of the limited time available for pixel stimulation and because of the limited laser power available. In practice, only about 30% of the PSL centres is stimulated to give rise to a PSL photon. Since these photons are emitted and scattered in all directions, only 50% of the PSL photons are emitted at the top side of the storage phosphor screen, where they can be detected by the detection system. The emitted PSL photons are guided towards the detector by a light guide. This light guide may consist of an array of optical fibres, that forms rectangular detection area above the storage phosphor screen and has a circular cross-section at the detector side. This type of light guide has a numerical aperture of only 30%, which means that only 1 out of 3 of the emitted PSL photons is guided to the detector. In between the light guide and the detector a filter is placed, which stops the stimulation light reflected by the storage phosphor screen and transmits the PSL light emitted by the screen. This filter also has a small absorption and reflection of PSL light and transmits only ca. 75% of the PSL photons. In computer radiography a photomultiplier is commonly used to transform the PSL signal into an electrical signal. At 440 nm the photomultiplier has a quantum efficiency of ca. 20%. This means that only 1 out of 5 PSL quanta that reach the photomultiplier are detected.
In summary, for 1,000 PSL centres that are created in the storage phosphor screen, only 1,000×0.3×0.5×0.3×0.75×0.2 or 6.75 PSL photons are detected. If it is required that every X-ray quantum gives rise to at least 1 detected PSL photon, therefore, the number of PSL centres created by an X-ray quantum should be sufficiently large. Or, conversely, the X-ray energy required to create a PSL-centre should be sufficiently small.
In mammography, a usual setting of the X-ray source is at 28 kVp. This leads to an X-ray spectrum, where the average energy of an X-ray quantum is of the order of 15 keV. For an X-ray quantum with this energy, in order to give rise to at least 1 detected PSL photon, the energy needed to create a PSL centre should be less than 15,000×6.75/1,000=100 eV.
Furtheron is well-known that fine detail visualisation, high-resolution high-contrast images are required for many X-ray medical imaging systems and particularly in mammography. The resolution of X-ray film/screen and digital mammography systems is currently limited to 20 line pairs/mm and 10 line pairs/mm, respectively. One of the key reasons for this limitation is associated with the phosphor particle size in the currently used X-ray screens. In particular, light scattering by the phosphor particles and their grain boundaries results in loss of spatial resolution and contrast in the image. In order to increase the resolution and contrast, scattering of the visible light must be decreased. Scattering can be decreased by reducing the phosphor particle size while maintaining the phosphor luminescence efficiency. Furthermore, the X-ray to light conversion efficiency, the quantum detection efficiency (e.g. the fraction of absorbed X-rays convertable to light emitted after stimulation) and the screen efficiency (e.g. the fraction of emitted light escaping from the screen after irradiation with stimulating rays) shouldn't be affected in a negative way by the reduction of the phosphor particle size. As a particular advantage the computed radiographic recording and reproducing techniques presented hereinbefore show a radiation image containing a large amount of information, obtainable with a markedly lower dose of radiation than in conventional radiography.
For clinical diagnosis and routine screening of asymptomatic female population, screen-film mammography today still represents the state-of-the-art for early detection of breast cancer. However, screen-film mammography has limitations which reduce its effectiveness. Because of the extremely low differentiation in radiation absorption densities in the breast tissue, image contrast is inherently low. Film noise and scatter radiation further reduce contrast making detection of microcalcifications difficult in the displayed image. So e.g. film gradient must be balanced against the need for wider latitude.
Digital radiography systems can be broadly categorized as primary digital and secondary digital systems. Primary digital systems imply direct conversion of the radiation incident on a sensor into usable electrical signals to form a digital image. Secondary digital systems, on the other hand, involve an intermediary step in the conversion of radiation to a digital image. For example, in digital fluoroscopy, image intensifiers are used for intermediary conversion of X-rays into a visible image which is then converted to a digital image using a video camera. Similarly, digital X-ray systems using photostimulated luminescence (PSL) plates, first store the virtual image as chemical energy. In a second step, the stored chemical energy is converted into electrical signals using a laser to scan the PSL plate to form a digital image.
Furthermore, various schemes using silicon photodiode arrays in scanning mode for digital radiography systems have been employed. However, these photodiode arrays require intermediate phosphor screens to convert X-rays into visible light, because of the steep fall-off in quantum efficiency (sensitivity) of the arrays at energies above 10 keV.
As is well-known a stimulable phosphor to be incorporated in the phosphor-incorporated area, i.e., a phosphor which absorbs not only a radiation having a wavelength of lower than 250 nm but also visible or ultraviolet light in the wavelength region of 250 to 400 nm, giving a stimulated emission of a wavelength in the range of 300 to 500 nm when it is irradiated with stimulating rays of a wavelength in the range of 400 to 900 nm, is preferably employed.
Examples of well-known, frequently used stimulable phosphors include divalent europium activated phosphors (e.g., BaFBr:Eu, BaFBrI:Eu) or cerium activated alkaline earth metal halide phosphors and cerium activated oxyhalide phosphors, as well as e.g. a phosphor having the formula of YLuSiO5:Ce,Zr.
In the present invention it is envisaged to use randomly one after another screen containing either, divalent europium activated alkali halide type phosphor screens, wherein said halide is at least one of chloride, bromide and iodide or a combination thereof divalent europium activated alkaline earth metal phosphor screens wherein said halide is at least one of fluoride, chloride, bromide and iodide or a combination thereof. Most preferred is random use of divalent europium activated CsX type phosphor screens, wherein said X represents Br or a combination of Br with at least one of Cl and I, as Br(Cl), Br(I) or Br(Cl, I) and bariumfluorohalide phosphor screens wherein the phosphor is of the (Ba,MII)FX′:Eu type, wherein MII is an alkaline earth metal and wherein X′ is Cl, Br and/or I.
Crystalline divalent europium activated alkali halide phosphor screens advantageously have CsBr:Eu2+ storage phosphor particles, in binderless layers in the form of cylinders (and even up to a needle-shaped form) wherein said cylinder has an average cross-section diameter in the range from 1 μm to 30 μm (more preferred: from 2 μm up to 15 μm), an average length, measured along the casing of said cylinder, in the range from 100 μm up to 1000 μm (more preferred: from 100 μm up to 500 μm) as has e.g. been described in EP-A 1 359 204. Such block-shaped, prismatic, cylindrical or needle-shaped phosphors, whether or not obtained after milling, are, in another embodiment, coated in a phosphor binder layer.
According to another embodiment of the present invention said stimulable phosphors are (Ba,MII)FX′:Eu type phosphors, wherein MII is an alkaline earth metal and wherein X′ is Cl, Br and/or I. In a preferred embodiment, said MII is Sr2+. Non-crystalline or amorphous europium activated alkaline earth metal halide phosphor screens advantageously have Ba(Sr)FBr:Eu2+ storage phosphor particles, dispersed in a binder medium in their corresponding storage phosphor layers, are advantageously used in the system of the present invention.
From the point of view of practical use, the differing stimulable phosphor screens or panels are desired, all of them, in order to give stimulated emission in the wavelength region of 300–500 nm when excited with stimulating rays in the wavelength region of 500–850 nm. This is particularly important when the detector is a photomultiplier having the highest quantum efficiency in the blue region. The stimulation light can only be filtered away when the wavelength of the emission light is quite different from the green or red stimulation light, i.e. that there is no or a only negligable overlap between stimulation radiation and stimulated emission radiation spectrum. In favor of customer-friendly handling or manutention in a medical radiographic environment, wherein a lot of phosphor plates or panels are exposed and read-out (processed) one after another, even if processed in a random order, it is recommended that detection of the blue photostimulated light proceeds with blue light transmitting filters for all screens or panels, without the need to change filters inbetween consecutive readings. Use of only one and same filter for all of the differing plates scanned in one and same scanning unit, in applications requiring optimum image quality as well as in applications requiring ordinary image quality, would be highly desired, in favor of cost reduction as only one scanner would be required needed for both types of plates.